Implantable pulse generator that generates spinal cord stimulation signals for a human body

ABSTRACT

An implantable pulse generator (IPG) that generates spinal cord stimulation signals for a human body has a programmable signal generator that can generate the signals based on stored signal parameters without any intervention from a processor that controls the overall operation of the IPG. While the signal generator is generating the signals the processor can be in a standby mode to substantially save battery power.

CROSS REFERENCE TO RELATED APPLICATIONS

The present application is a continuation-in-part of U.S. patentapplication Ser. No. 14/805,607, filed Jul. 22, 2015 (now issued as U.S.Pat. No. 9,517,343), which is a continuation-in-part of U.S. patentapplication Ser. No. 14/213,186, filed Mar. 14, 2014 (now issued as U.S.Pat. No. 9,492,665), which claims priority to U.S. ProvisionalApplication No. 61/792,654, filed Mar. 15, 2013, entitled “SPINAL CORDSTIMULATOR SYSTEM,” all of which are herein incorporated by reference intheir entirety.

TECHNICAL FIELD

This disclosure relates to stimulators using electrical pulses in amedical context, and more particularly, applying electrical stimulationsignals to the spinal cord to control pain.

BACKGROUND

A Spinal Cord Stimulator (SCS) is used to exert pulsed electricalsignals to the spinal cord to control chronic pain. Spinal cordstimulation, in its simplest form, comprises stimulating electrodesimplanted in the epidural space, an implantable pulse generatorimplanted in the lower abdominal area or gluteal region, conductingwires connecting the electrodes to the electrical pulse generator, anelectrical pulse generator remote control, and an electrical pulsegenerator charger. Spinal cord stimulation has notable analgesicproperties and, at the present, is used mostly in the treatment offailed back surgery syndrome, complex regional pain syndrome andrefractory pain due to ischemia.

Electrotherapy of pain by neurostimulation began shortly after Melzackand Wall proposed the gate control theory in 1965. This theory proposedthat nerves carrying painful peripheral stimuli and nerves carryingtouch and vibratory sensation both terminate in the dorsal horn (thegate) of the spinal cord. It was hypothesized that input to the dorsalhorn of the spinal cord could be manipulated to “close the gate” to thenerves. As an application of the gate control theory, Shealy et al.implanted the first spinal cord stimulator device directly on the dorsalcolumn for the treatment of chronic pain in 1971.

Spinal cord stimulation does not eliminate pain. The electrical impulsesfrom the stimulator override the pain messages so that the patient doesnot feel the pain intensely. In essence, the stimulator masks the pain.A trial implantation is performed before implanting the permanentstimulator. The physician first implants a trial stimulator through theskin (percutaneously) to perform stimulations as a trial run. Because apercutaneous trial stimulator tends to move from its original location,it is considered temporary. If the trial is successful, the physiciancan then implant a permanent stimulator. The permanent stimulator isimplanted under the skin of the abdomen with the leads inserted underthe skin and subcutaneously fed to and inserted into the spinal canal.This placement of the stimulator in the abdomen is a more stable,effective location. The leads, which consist of an array of electrodes,can be percutaneous type or paddle type. Percutaneous electrodes areeasier to insert in comparison with paddle type, which are inserted viaincision over spinal cord and laminectomy.

There are a number of problems that exist in currently availableimplantable pulse generators that limit the full benefits of dorsalcolumn stimulation from an effectiveness and patient user friendlyperspective.

One problem is that the circuits in the current generators consume toomuch power. This requires frequent recharging, making it veryinconvenient for patients. Another problem is that the currentgenerators are limited in concurrently generating different stimulationpatterns to treat different parts of the body simultaneously.Accordingly, when patients have varying degrees of pain in differentparts of the body, it is difficult, if not impossible, to effectivelytreat all area of pain.

Therefore, it would be desirable to provide a system and method forgenerating stimulation patterns which resolve the problems discussedabove.

SUMMARY OF THE DISCLOSURE

According to one aspect of the present invention, there is provided animplantable pulse generator (IPG) that generates spinal cord stimulationsignals for a human body has a programmable signal generator that cangenerate the signals based on stored signal parameters without anyintervention from a processor that controls the overall operation of theIPG. While the signal generator is generating the signals the processorcan be in a standby mode to substantially save battery power.

According to another aspect of the present invention, there is providedan implantable pulse generator (IPG) that generates spinal cordstimulation signals for a human body has control registers that storestimulation signal parameters for stimulation channels with each channelcapable of being associated with at least two electrodes andrepresenting a particular stimulation signal pattern for the associatedelectrodes. An arbitrator continuously receives timing signalsrepresenting the stimulation signal patterns and selects one channelamong the many channels as an active treatment channel in order to avoidtwo channels from being activated at the same time. The arbitratorprovides flexibility in programming different pulse parameters formultiple stimulation channels without the possibility of overloading thepower supply that generates the stimulation signal patterns.

According to another aspect of the present invention, there is providedan implantable pulse generator (IPG) that generates spinal cordstimulation signals for a human body, which includes a timing generatorand high frequency generator. The timing generator generates timingsignals that represent stimulation signals for multiple channels. Thehigh frequency generator determines whether to modulate the timingsignals and modulates them at a burst frequency according to storedburst parameters if the decision is yes. As such, the IPG provides theability to generate both the low frequency and high frequencystimulation signals in different channels according to user programming.

According to another aspect of the present invention, there is providedan implantable pulse generator (IPG) that generates spinal cordstimulation signals for a human body, which includes a timing generatorand high frequency generator. The timing generator generates timingsignals that represent stimulation signals for multiple channels. Thehigh frequency generator determines whether to modulate the timingsignals and modulates them at a burst frequency according to storedburst parameters if the decision is yes. The high frequency generatorcan also independently control the pulse frequency of each channelaccording to the stored parameters. As such, the IPG provides theability to generate both the low frequency and high frequencystimulation signals at different frequencies in different channelsaccording to user programming in order to provide maximum flexibility intreatment.

According to another aspect of the present invention, there is providedan implantable pulse generator (IPG) that generates spinal cordstimulation signals for a human body, which includes an electrode driverfor each electrode, which adjusts the amplitude of the timing signalsand output an output current corresponding to the adjusted signals fortransmission to the associated electrode so as to enable independentamplitude control of the stimulation signals for each stimulationpattern channel.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts various components that can be included in a spinal cordstimulation system, according to an embodiment, during trial andpermanent implantation.

FIG. 2 depicts an exploded view of an implantable pulse generator (IPG)assembly, according to an embodiment.

FIG. 3 depicts a feedthrough assembly of an implantable pulse generator(IPG) assembly, according to an embodiment.

FIG. 4 depicts a lead contact system of an implantable pulse generator(IPG) assembly, according to an embodiment.

FIG. 5 depicts a lead contact assembly of an implantable pulse generator(IPG) assembly, according to an embodiment.

FIG. 6 depicts a head unit assembly of an implantable pulse generator(IPG) assembly, according to an embodiment.

FIG. 7 depicts an RF antenna of an implantable pulse generator (IPG)assembly, according to an embodiment.

FIG. 8 depicts a percutaneous lead, according to an embodiment.

FIG. 9 depicts a paddle lead, according to an embodiment.

FIG. 10 depicts a lead extension, according to an embodiment.

FIG. 11 depicts a lead splitter, according to an embodiment.

FIG. 12 depicts a sleeve anchor, according to an embodiment.

FIG. 13 depicts a mechanical locking anchor, according to an embodiment.

FIG. 14 illustrates communication via a wireless dongle with atablet/clinician programmer and smartphone/mobile/patient programmerduring trial and/or permanent implantation, according to an embodiment.

FIG. 15 depicts a Tuohy needle, according to an embodiment.

FIG. 16 depicts a stylet, according to an embodiment.

FIG. 17 depicts a passing elevator, according to an embodiment.

FIG. 18 depicts a tunneling tool, according to an embodiment.

FIG. 19 depicts a torque wrench, according to an embodiment.

FIG. 20 is a function block diagram of some components in an implantablepulse generator according to an embodiment.

FIG. 21 is a functional block diagram of the signal generator of FIG.20.

FIG. 22 is a functional block diagram of the high frequency generator ofFIG. 21.

FIG. 23 is a functional block diagram of the electrode driver of FIG.21.

FIG. 24 is a functional illustration of two of the current drivers ofFIG. 23.

FIG. 25 illustrates a grouping of electrodes for different channelsaccording to an embodiment of the present invention.

FIG. 26 shows exemplary electrode waveforms illustrating an asymmetricalpulse pattern between a positive and negative pulse according to anembodiment of the present invention.

FIG. 27 shows exemplary electrode waveforms for two arbitratedstimulation channels according to an embodiment of the presentinvention.

DETAILED DESCRIPTION

Implantable Pulse Generator (IPG)

FIG. 1 illustrates various components that can be included in a SCSsystem for the trial and the permanent installation periods. The spinalcord stimulator (SCS) 100 is an implantable device used to deliverelectrical pulse therapy to the spinal cord in order to treat chronicpain. The implantable components of the system consist of an ImplantablePulse Generator (IPG) 102 and a multitude of stimulation electrodes 130.In some embodiments, the IPG 102 is implanted subcutaneously, no morethan 30 mm deep in an area that is comfortable for the patient while thestimulation electrodes 130 are implanted directly in the epidural space.In other embodiments, the IPG 102 is implanted no more than 20 mm, 25mm, 35 mm or 40 mm. The electrodes 130 are wired to the IPG 102 vialeads 140, 141 which keep the stimulation pulses isolated from eachother in order to deliver the correct therapy to each individualelectrode 130.

The therapy delivered consists of electrical pulses with controlledcurrent amplitude ranging from +12.7 to −12.7 mA (current range 0-25.4mA). In other embodiments, the amplitude can range from +13.7 to −13.7mA, or +14.7 to −14.7 mA. These pulses can be programmed in both lengthand frequency from 100 to 20000 and 0.5 Hz to 1200 Hz. At any givenmoment, the sum of the currents sourced from the anodic electrodes 130can equal the sum of the currents sunk by the cathodic electrodes 130.In addition, each individual pulse is bi-phasic, meaning that once theinitial pulse finishes another pulse of opposite amplitude is generatedafter a set holdoff period. The electrodes 130 may be grouped intostimulation sets in order to deliver the pulses over a wider area or totarget specific areas, but the sum of the currents being sourced at anyone given time may not exceed 20 mA in accordance with some embodiments.In other embodiments, the sum of the currents being sourced at any onegiven time may not exceed 15 mA, 25 mA, 30 mA or greater. A user canalso program different stimulation sets (e.g., eight, ten, twelve ormore) with different parameters in order to target different areas withdifferent therapies.

FIG. 2 depicts an exploded view of an IPG 102. The IPG 102 consists oftwo major active components 104, 106, a battery 108, antenna 110, somesupport circuitry, and a multitude of output capacitors 112. The firstof the major active components is the microcontroller transceiver 104.It is responsible for receiving, decoding, and execution both commandsand requests from the external remote. If necessary it passes thesecommands or requests onto the second major component, the ASIC 106. TheASIC 106 receives the digital data from the microcontroller 104 andperforms the entire signal processing to generate the signals necessaryfor stimulation. These signals are then passed onto the stimulationelectrodes 130 in the epidural space.

The ASIC 106 is comprised of a digital section and an analog section. Insome embodiments, the digital section is divided into multiple sectionsincluding; Timing Generators, Arbitration Control, Pulse BurstConditioner, and Electrode Logic. The analog section receives theincoming pulses from the digital section and amplifies them in order todeliver the correct therapy. There are also a multitude of digitalregister memory elements that each section utilizes, both digital andanalog.

The digital elements in the ASIC 106 are all made up of standard subsetsof digital logic including logic gates, timers, counters, registers,comparators, flip-flops, and decoders. These elements are ideal forprocessing the stimulation pulses as all of them can function extremelyfast—orders of magnitudes faster than the required pulse width. In someembodiments, the elements all function at one single voltage, usually5.0, 3.3, 2.5, or 1.8 volts.

The timing generators are the base of each of the stimulation sets. Itgenerates the actual rising and falling edge triggers for each phase ofthe bi-phasic pulse. It accomplishes this by taking the incoming clockthat is fed from the microcontroller 104 and feeding it into a counter.For the purpose of this discussion, assume the counter simply countsthese rising clock edges infinitely. The output of the counter is fedinto six different comparators. The comparators other input is connectedto specific registers that are programmed by the microcontroller 104.When the count equals the value stored in the register, the comparatorasserts a positive signal.

The first comparator is connected to the SET signal of a SR flip flop.In some embodiments, the SR flip flop stays positive until the RESETsignal is asserted, which the second comparator is connected to. Theoutput of the SR flip flop is the first phase of the bi-phasic pulse.Its rising & falling edges are values stored in the registers andprogrammed by the microcontroller 104. The third and fourth comparators& registers work in exactly the same way to produce the second phase ofthe bi-phasic pulse using the second SR flip flop.

The fifth comparator is connected the RESET of the final SR-Flip flop inthe timing generator. This flip flop is SET by the first comparator,which is the rising edge of the first pulse. The RESET is then triggeredby the value the microprocessor programmed into the register connectedto the comparator. This allows for a ‘holdoff’ period after the fallingedge of the second pulse. The output of this third SR flip flop can bethought of as an envelope of the biphasic pulses indicating when thisparticular timing generator is active.

The final comparator of the system is once again connected to a registerthat stores the frequency values from the microprocessor. Essentiallywhen the count reaches this value it triggers the comparator which isfed back to the counter to reset it to zero and beginning the entirepulse generation cycle again. The ASIC 106 may contain many of thesetiming generators as each can control anywhere from two to all of theelectrodes 130 connected to the IPG 102 at a time. However, when thereis more than one timing generator and multiple channels have beenactively programmed then there needs to be a mechanism for suppressing asecond channel from turning on when another is already active.

The next circuit block contained in the IPG 102 is the arbitrator. Thearbitrator functions by looking at each of the timing generators'envelope signals and makes sure only one can be active at a time. If asecond tries to activate then the arbitrator suppresses that signal.

The arbitrator accomplishes this by bringing each of the channelenvelope signals into a rising edge detection circuit. Once one istriggered it is fed into the SET pin of an SR flip flop. The output ofthis SR-flip flop is fed into all of the other rising edge detectors inorder to suppress them from triggering. The channel envelope signal isalso fed into a falling-edge detector which is then fed into the RESETof the same SR flip flop. The output of the SR flip flops are thenconnected to switches whose outputs are all tied together that turnon/off that channels particular biphasic pulse train. Therefore, theoutput of this circuit element is a single bi-phasic pulse train and asignal designating which timing generator that particular pulse train issourced from. Essentially, the circuit looks for a channel to go active.Once it finds one it suppresses all others until that channel becomesinactive.

The next section of the circuit works very similarly to the timinggenerators to create a high speed burst pulse train that is thencombined with the stimulation pulse train to create a bursted bi-phasicpulse train if desired.

It accomplishes this by taking the incoming clock that is fed from themicrocontroller 104 and feeding it into a counter. The counter can countthese rising clock edges infinitely. The counter is only active during asingle phase of the bi-phasic signal and begins counting as soon as therising edge is detected. The output of the counter is fed into acomparator, along with a microcontroller-programmed register, whoseoutput is connected to the reset pin on the counter. Therefore, thiscounter will simply count to a programmed value and reset. Thisprogrammed value is the burst frequency.

The output of the comparator is then fed into an edge detection circuitand then a flip flop that combines it with the actual stimulation pulsetrain to create a single phase bursted stimulation pulse. The entirecircuit is duplicated for the second phase of the signal resulting inthe desired bursted bi-phasic pulse train. The stimulation signal is nowhanded over to the electrode logic stage.

The electrode logic conditions and directs the bi-phasic signals to theanalog section of the ASIC 106. At this point, the bi-phasic signalscontain all of the pertinent timing information, but none of therequired amplitude information. The incoming signals include thebi-phasic pulse train and another signal designating which timinggenerator the current active train came from. Each electrode logic cellhas a register for each timing generator that stores this particularelectrode's 130 amplitude values for that timing generator. Theelectrode logic cell uses the designation signal to determine whichregister to pull the amplitude values from, e.g. if the third timinggenerator is passed through the arbitration circuit then the electrodelogic would read the value from the third register.

Once the value is pulled from the register, it goes through a series oflogic gates. The gates first determine that the electrode 130 should beactive. If not, no further action is taken and the analog section of theelectrode output is not activated, thereby saving precious battery 108power. Next, a determination is made if the particular electrode 130 isan anode or cathode. If the electrode is deemed to be an anode, theelectrode logic passes the amplitude information and the biphasic signalto the positive current (digital to analog converter) DAC in the analogsection of the ASIC 106. If the electrode is deemed to be a cathode, theelectrode logic passes the amplitude information and the biphasic signalto the negative current DAC in the analog section of the ASIC 106. Insome embodiments, the electrode logic circuit makes these decisions foreach phase of the bi-phasic signal as every electrode 130 will switchbetween being an anode and a cathode.

The analog elements in the ASIC 106 are uniquely designed in order toproduce the desired signals. The basis of analog IC design is the fieldeffect transistor (FET) and the type of high current multiple outputdesign required in SCS 100 means that the bulk of the silicon in theASIC 106 will be dedicated to the analog section.

The signals from the electrode output are fed into each current DAC whenthat specific electrode 130 should be activated. Each electrode 130 hasa positive and a negative current DAC, triggered by the electrode logicand both are never active at the same time. The job of each current DACis, when activated, to take the digital value representing a stimulationcurrent amplitude and produce an analog representation of this value tobe fed into the output stage. This circuit forms half of the barrierbetween the digital and analog sections of the ASIC 106.

The digital section of the ASIC 106 is built upon a technology that onlyallows small voltages to exist. In moving to the analog section, theoutput of the current DAC (which is a low level analog signal) can beamplified to a higher voltage for use in the analog section. The circuitthat performs this task is called a power level shifter. Because thiscircuit is built upon two different manufacturing technologies andrequires high precision analog circuits built upon a digital base, itcan be difficult to implement.

Once the voltages have been converted for usage in the analog portion ofthe ASIC 106 the voltages are passed on to the output current stages.There are two current sources per electrode output. In some embodiments,one will source a positive current and one will sink a negative current,but both will not be active simultaneously. The current sourcesthemselves are made up of analog elements similar to a Howland currentsource. There is an input stage, and an amplification stage withfeedback through a sensing component to maintain the constant current.The input stage takes the analog voltage values from the power levelshifter and produces an output pulse designated for the amplifier. Theamplifier then creates the pulses of varying voltages but constantcurrent flow. The sources are capable of sourcing or sinking up to 12.7mA at 0.1 mA resolution into a load of up to 1.2 k Ohms. This translatesinto range of 15 volts, which will vary depending on the load in orderto keep the current constant.

The microcontroller 104 to ASIC 106 interface is designed to be assimple as possible with minimal bus ‘chatter’ in order to save battery108 life. The ASIC 106 can be a collection of registers programmed via astandard I²C or SPI bus. Since the ASIC 106 is handling all the powermanagement, there will also be a power good (PG) line between the twochips 104, 106 in order to let the microcontroller 104 know when it issafe to power up. The ASIC 106 will also need to use a pin on themicrocontroller 104 in order to generate a hardware interrupt in caseanything goes awry in the ASIC 106. The final connection is the timebase for all of the stimulation circuitry. In some embodiments, the ASIC106 will utilize two clocks, one for its internal digital circuitrywhich will be fed directly from the microcontroller 104 clock output,and one to base all stimulation off of which will need to be synthesizedby the microcontroller 104 and fed to the ASIC 106. All commands andrequests to the ASIC 106 will be made over the I²C or SPI bus and willinvolve reading a register address or writing to a register. Even whenthe ASIC 106 generates a hardware interrupt, it will be theresponsibility of the microcontroller 104 to poll the ASIC 106 anddetermine the cause of the interrupt.

The wireless interface is based upon the FCCs MedRadio standardoperating in the 402-405 MHz range utilizing up to 10 channels fortelemetry. The protocol implemented is chosen to minimize transmissionand maximize battery 108 life. All processing can take place on the userremote/programmer and the only data transmitted is exactly what will beused in the microcontroller 104 to ASIC 106 bus. That is, all of thewireless packets will contain necessary overhead information along withonly a register address, data to store in the register, and a commandbyte instructing the microcontroller 104 what to do with the data. Theoverhead section of the wireless protocol will contain synchronizationbits, start bytes, an address which is synchronized with the IPG's 102serial number, and a CRC byte to assure proper transmission. The packetlength is kept as small as possible in order to maintain battery 108life. Since the IPG 102 cannot listen for packets all the time due tobattery 108 life, it cycles on for a duty cycle of less than 0.05% ofthe time. This time value can be kept small as long as the data packetsare also small. The user commands needed to run the system are executedby the entire system using flows.

The IPG 102 uses an implantable grade Li ion battery 108 with 215 mAHrwith zero volt technology. The voltage of the battery 108 at fullcapacity is 4.1 V and it supplies current only until it is drained up to3.3 V which is considered as 100% discharged. The remaining capacity ofthe battery 108 can be estimated at any time by measuring the voltageacross the terminals. The maximum charge rate is 107.5 mA. A ConstantCurrent, Constant Voltage (CCCV) type of regulation can be applied forfaster charging of the battery 108.

The internal secondary coil 109 is made up of 30 turns of 30 AWG coppermagnet wires. The ID, OD, and the thickness of the coil are 30, 32, and2 mm, respectively. Inductance L2 is measured to be 58 uH, a 80 nFcapacitor is connected to it to make a series resonance tank at 74 kHzfrequency. In the art of induction charging, two types of rectifiers areconsidered to convert the induced AC into usable DC, either a bridgefull wave rectifier or a voltage doubler full wave rectifier. To obtaina higher voltage, the voltage double full wave rectifier is used in thisapplication. The rectifier is built with high speed Schottky diodes toimprove its function at high frequencies of the order 100 kHz. A Zenerdiode and also a 5V voltage regulator are used for regulation. Thiscircuit will be able to induce AC voltage, rectify to DC, regulate to 5Vand supply 100 mA current to power management IC that charges theinternal battery 108 by CCCV regulation.

The regulated 5V 100 mA output from the resonance tank is fed to, forexample, a Power Management Integrated Circuit (PMIC) MCP73843. Thisparticular chip was specially designed by Microchip to charge a Li ionbattery 108 to 4.1 V by CCCV regulation. The fast charge current can beregulated by changing a resistor; it is set to threshold current of 96mA in the example circuit. The chip charges the battery 108 to 4.1V aslong as the current received is more than 96 mA. However, if the supplycurrent drops below 96 mA, it stops to charge the battery 108 until thesupply is higher than 96 again. For various practical reasons, if thedistance between the coils increases, the internal secondary coil 109receives lesser current than the regulated value, and instead ofcharging the battery 108 slowly, it pauses the charging completely untilit receives more than 96 mA. It is understood to those with skill in theart that other power management chips can be used and the powermanagement chip is not limited to the PMIC MCP738432 chip.

All the functions of the IPG 102 are controlled from outside using ahand held remote controller specially designed for this device. Alongwith the remote control, an additional control is desirable to operatethe IPG 102 if the remote control was lost or damaged. For this purposea Hall effect based magnet switch was incorporated to either turn ON orturn OFF the IPG 102 using an external piece of magnet. Magnet switchacts as a master control for the IPG 102 to turn on or off. A south poleof sufficient strength turns the output on and a north pole ofsufficient strength turns the output off. The output is latched so thatthe switch continues to hold the state even after the magnet is removedfrom its vicinity.

The IPG 102 is an active medical implant that generates an electricalsignal that stimulates the spinal cord. The signal is carried through astimulation lead 140 that plugs directly into the IPG 102. The IPG 102recharges wirelessly through an induction coil 109, and communicates viaRF radio antenna 110 to change stimulation parameters. In someembodiments, the IPG 102 is implanted up to 2 cm, 3 cm, 4 cm or 5 cmbelow the surface of the skin and can be fixed to the fascia by passingtwo sutures through holes in the epoxy header 114. The leads 140 areelectrically connected to the IPG 102 through a lead contact system 116,a cylindrical spring-based contact system with inter-contact siliconeseals. The leads 140 are secured to the IPG 102 with a set screw 117that actuates within locking housing 118. Set screw compression on thelead's 140 fixation contact can be governed by a disposable torquewrench. The wireless recharging is achieved by aligning the exteriorinduction coil on the charger with the internal induction coil 109within the IPG 102. The RF antenna within the remote's dongle 200communicates with the RF antenna 110 in the IPG's 102 epoxy header 114.FIG. 2 illustrates an exploded view of the IPG 102 assembly.

The IPG 102 is an assembly of a hermetic titanium (6Al-4V) casing 120which houses the battery 108, circuitry 104, 106, and charging coil 109.The IPG 102 further includes an epoxy header 114 (see FIG. 6), whichhouses the lead contact assembly 116, locking housing 118, and RFantenna 110 (see FIGS. 6 and 7). The internal electronics are connectedto the components within the epoxy head through a hermetic feedthrough122, as shown in FIG. 3. The feedthrough 122 is a titanium (6Al-4V)flange with an alumina window and gold trimming. Within the aluminawindow are thirty-four platinum-iridium (90-10) pins that interfaceinternally with a direct solder to the circuit board, and externallywith a series of platinum iridium wires laser-welded to the antenna 110and lead contacts 126. The IPG 102 interfaces with 32 electricalcontacts 126, which are arranged in four rows of eight contacts 126.Thirty two of the feedthrough's 122 pins 124 interface with the contacts126, while two interface with the antenna 110, one to the ground planeand one to the antenna 110 feed.

FIGS. 4 and 5 depict a lead contact system 115 and assembly 116,respectively. The lead contacts 126 consist of an MP35N housing 128 witha platinum-iridium 90-10 spring 129. Each contact 126 is separated by asilicone seal 127. At the proximal end of each stack of 8 contacts 126is a titanium (6Al-4V) cap 125 which acts as a stop for the lead 140. Atthe distal end is a titanium (6Al-4V) set screw 119 and block 118 forlead fixation. At the lead entrance point is a silicone tube 123 whichprovides strain relief as the lead 140 exits the head unit 114, andabove the set screw 119 another silicone tube 131 with a small internalcanal allows the torque wrench to enter but does not allow the set screw119 to back out. In addition to the contacts 126 and antenna 110, theheader 114 also contains a radiopaque titanium (6Al-4V) tag 132 whichallows for identification of the device under fluoroscopy. The overmoldof the header 114 is Epotek 301, a two-part, biocompatible epoxy. FIGS.4, 5, 6, and 7 depict illustrations of lead contact system 115, leadcontact assembly 116, head unit assembly 114, and RF antenna 110,respectively.

Internal to the titanium (6Al-4V) case 120 are the circuit board 105,battery 108, charging coil 109, and internal plastic support frame. Thecircuit board 105 can be a multi-layered FR-4 board with copper tracesand solder mask coating. Non-solder masked areas of the board can beelectroless nickel immersion gold. The implantable battery 108, allsurface mount components, ASIC 106, microcontroller 104, charging coil109, and feedthrough 122 will be soldered to the circuit board 105. Theplastic frame, made of either polycarbonate or ABS, will maintain thebattery's 108 position and provide a snug fit between the circuitry 105and case 120 to prevent movement. The charging coil 109 is a woundcoated copper.

Leads

The percutaneous stimulation leads 140, as depicted in FIG. 8, are afully implantable electrical medical accessory to be used in conjunctionwith the implantable SCS 100. The primary function of the lead is tocarry electrical signals from the IPG 102 to the target stimulation areaon the spinal cord. Percutaneous stimulation leads 140 providecircumferential stimulation. The percutaneous stimulation leads 140provide a robust, flexible, and bio-compatible electric connectionbetween the IPG 102 and stimulation area. The leads 140 are surgicallyimplanted through a spinal needle, or epidural needle, and are driventhrough the spinal canal using a steering stylet that passes through thecenter of the lead 140. The leads 140 are secured mechanically to thepatient using either an anchor or a suture passed through tissue andtied around the body of the lead 140. The leads 140 are secured at theproximal end with a set-screw 119 on the IPG 102 which applies radialpressure to a blank contact on the distal end of the proximal contacts.

The percutaneous stimulation leads 140 consist of a combination ofimplantable materials. Stimulation electrodes 130 at the distal end andelectrical contacts at the proximal end are made of a 90-10platinum-iridium alloy. This alloy is utilized for its bio-compatibilityand electrical conductivity. The electrodes 130 are geometricallycylindrical. The polymeric body of the lead 140 is polyurethane, chosenfor its bio-compatibility, flexibility, and high lubricity to decreasefriction while being passed through tissue. The polyurethane tubing hasa multi-lumen cross section, with one center lumen 142 and eight outerlumens 144. The center lumen 142 acts as a canal to contain the steeringstylet during implantation, while the outer lumens 144 provideelectrical and mechanical separation between the wires 146 that carrystimulation from the proximal contacts to distal electrodes 130. Thesewires 146 are a bundle of MP35N strands with a 28% silver core. Thewires 146 are individually coated with ethylene tetrafluoroethylene(ETFE), to provide an additional non-conductive barrier. The wires 146are laser welded to the contacts and electrodes 130, creating anelectrical connection between respective contacts on the proximal anddistal ends. The leads 140 employ a platinum-iridium plug 148, moldedinto the distal tip of the center lumen 142 to prevent the tip of thesteering stylet from puncturing the distal tip of the lead 140. Leads140 are available in a variety of 4 and 8 electrode 130 configurations.These leads 140 have 4 and 8 proximal contacts (+1 fixation contact),respectively. Configurations vary by electrode 130 number, electrode 130spacing, electrode 130 length, and overall lead 140 length.

The paddle stimulation leads 141, as depicted in FIG. 9, are a fullyimplantable electrical medical accessory to be used in conjunction withthe implantable SCS 100. The primary function of the paddle lead 141 isto carry electrical signals from the IPG 102 to the target stimulationarea on the spinal cord. In some embodiments, the paddle leads 141provide uni-direction stimulation across a 2-dimensional array ofelectrodes 130, allowing for greater precision in targeting stimulationzones. The paddle stimulation leads 141 provide a robust, flexible, andbio-compatible electric connection between the IPG 102 and stimulationarea. The leads 141 are surgically implanted through a small incision,usually in conjunction with a laminotomy or laminectomy, and arepositioned using forceps or a similar surgical tool. The leads 141 aresecured mechanically to the patient using either an anchor or a suturepassed through tissue and tied around the body of the lead 141. Theleads 141 are secured at the proximal end with a set-screw on the IPG102 which applies radial pressure to a fixation contact on the distalend of the proximal contacts.

The paddle stimulation leads 141 consist of a combination of implantablematerials. Stimulation electrodes 130 at the distal end and electricalcontacts at the proximal end are made of a 90-10 platinum-iridium alloyutilized for its bio-compatibility and electrical conductivity. Thepolymeric body of the lead 141 is polyurethane, chosen for itsbio-compatibility, flexibility, and high lubricity to decrease frictionwhile being passed through tissue. The polyurethane tubing has amulti-lumen cross section, with one center lumen 142 and eight outerlumens 144. The center lumen 142 acts as a canal to contain the steeringstylet during implantation, while the outer lumens 144 provideelectrical and mechanical separation between the wires 146 that carrystimulation from the proximal contacts to distal electrodes 130. Thesewires 146 are a bundle of MP35N strands with a 28% silver core. Thewires 146 are individually coated with ethylene tetrafluoroethylene(ETFE), to provide an additional non-conductive barrier. At the distaltip of the paddle leads 141 is a 2-dimensional array of flat rectangularelectrodes 130 molded into a flat silicone body 149. In an embodiment,one side of the rectangular electrodes 130 is exposed, providinguni-directional stimulation. The wires 146 are laser welded to thecontacts and electrodes 130, creating an electrical connection betweenrespective contacts on the proximal and distal ends. Also molded intothe distal silicone paddle is a polyester mesh 147 adding stability tothe molded body 149 while improving aesthetics by covering wire 146routing. The number of individual 8-contact leads 141 used for eachpaddle 141 is governed by the number of electrodes 130. In someembodiments, electrodes 130 per paddle range from 8 to 40, split intobetween one and four proximal lead 141 ends. In some embodiments, thereare 32 electrodes 130 per paddle. Each proximal lead 141 has 8 contacts(+1 fixation contact). Configurations vary by electrode 130 number,electrode 130 spacing, electrode length, and overall lead length.

The lead extensions 150, as depicted in FIG. 10, are a fully implantableelectrical medical accessory to be used in conjunction with theimplantable SCS 100 and either percutaneous 140 or paddle 141 leads. Theprimary function of the lead extension 150 is to increase the overalllength of the lead 140, 141 by carrying electrical signals from the IPG102 to the proximal end of the stimulation lead 140, 141. This extendsthe overall range of the lead 140, 141 in cases where the length of theprovided leads 140, 141 are insufficient. The lead extensions 150provide a robust, flexible, and bio-compatible electric connectionbetween the IPG 102 and proximal end of the stimulation lead 140, 141.The extensions 150 may be secured mechanically to the patient usingeither an anchor or a suture passed through tissue and tied around thebody of the extension 150. Extensions 150 are secured at the proximalend with a set-screw 119 on the IPG 102 which applies radial pressure toa fixation contact on the distal end of the proximal contacts of theextension 150. The stimulation lead 140, 141 is secured to the extension150 in a similar fashion, using a set screw 152 inside the molded tip ofextension 150 to apply a radial pressure to the fixation contact at theproximal end of the stimulation lead 140, 141.

The lead extension 150 consists of a combination of implantablematerials. In some embodiments, at the distal tip of the extension 150is a 1×8 array of implantable electrical contacts 154, each consistingof MP35 housing 128 and 90-10 platinum-iridium spring. A silicone seal127 separates each of the housings 128. At the proximal end of thecontacts is a titanium (6Al4V) cap which acts as a stop for the lead,and at the distal tip, a titanium (6Al4V) block and set screw 152 forlead fixation. The electrical contacts at the proximal end are made of a90-10 platinum-iridium alloy utilized for its bio-compatibility andelectrical conductivity. The polymeric body 156 of the lead 150 ispolyurethane, chosen for its bio-compatibility, flexibility, and highlubricity to decrease friction while being passed through tissue. Thepolyurethane tubing 158 has a multi-lumen cross section, with one centerlumen 142 and eight outer lumens 144. The center lumen 142 acts as acanal to contain the steering stylet during implantation, while theouter lumens 144 provide electrical and mechanical separation betweenthe wires 146 that carry stimulation from the proximal contacts todistal electrodes. These wires 146 are a bundle of MP35N strands with a28% silver core. The wires 146 are individually coated with ethylenetetrafluoroethylene (ETFE), to provide an additional non-conductivebarrier. Each lead extension 150 has 8 proximal cylindrical contacts (+1fixation contact).

The lead splitter 160, as depicted in FIG. 11, is a fully implantableelectrical medical accessory which is used in conjunction with the SCS100 and typically a pair of 4-contact percutaneous leads 140. Theprimary function of the lead splitter 160 is to split a single lead 140of eight contacts into a pair of 4 contact leads 140. The splitter 160carries electrical signals from the IPG 102 to the proximal end of two4-contact percutaneous stimulation leads 140. This allows the surgeonaccess to more stimulation areas by increasing the number of stimulationleads 140 available. The lead splitter 160 provides a robust, flexible,and bio-compatible electrical connection between the IPG 102 andproximal ends of the stimulation leads 140. The splitters 160 may besecured mechanically to the patient using either an anchor or a suturepassed through tissue and tied around the body of the splitter 160.Splitters 160 are secured at the proximal end with a set-screw 119 onthe IPG 102 which applies radial pressure to a fixation contact on thedistal end of the proximal contacts of the splitter 160. The stimulationleads 140 are secured to the splitter 160 in a similar fashion, using apair of set screws inside the molded tip of splitter 160 to apply aradial pressure to the fixation contact at the proximal end of eachstimulation lead 140.

The lead splitter 160 consists of a combination of implantablematerials. At the distal tip of the splitter 160 is a 2×4 array ofimplantable electrical contacts 162, with each contact 162 consisting ofMP35 housing 128 and 90-10 platinum-iridium spring. A silicone seal 127separates each of the housings 128. At the proximal end of each row ofcontacts 162 is a titanium (6Al4V) cap which acts as a stop for thelead, and at the distal tip, a titanium (6Al4V) block and set screw forlead fixation. The electrical contacts at the proximal end of thesplitter 160 are made of a 90-10 platinum-iridium alloy utilized for itsbio-compatibility and electrical conductivity. The polymeric body 164 ofthe lead 160 is polyurethane, chosen for its bio-compatibility,flexibility, and high lubricity to decrease friction while being passedthrough tissue. The polyurethane tubing 166 has a multi-lumen crosssection, with one center lumen 142 and eight outer lumens 144. Thecenter lumen 142 acts as a canal to contain the steering stylet duringimplantation, while the outer lumens 144 provide electrical andmechanical separation between the wires 146 that carry stimulation fromthe proximal contacts to distal electrodes 130. These wires 146 are abundle of MP35N strands with a 28% silver core. The wires 146 areindividually coated with ethylene tetrafluoroethylene (ETFE), to providean additional non-conductive barrier. Each lead splitter 160 has 8proximal contacts (+1 fixation contact), and 2 rows of 4 contacts 162 atthe distal end.

Anchors

The lead anchor 170, as depicted in FIGS. 12 and 13, is a fullyimplantable electrical medical accessory which is used in conjunctionwith both percutaneous 140 and paddle 141 stimulation leads. The primaryfunction of the lead anchor 170 is to prevent migration of the distaltip of the lead 140, 141 by mechanically locking the lead 140, 141 tothe tissue. There are currently two types of anchors 170, a simplesleeve 171, depicted in FIG. 12, and a locking mechanism 172, depictedin FIG. 13, and each has a slightly different interface. For the simplesleeve type anchor 171, the lead 140, 141 is passed through the centerthru-hole 174 of the anchor 171, and then a suture is passed around theoutside of the anchor 171 and tightened to secure the lead 140, 141within the anchor 171. The anchor 171 can then be sutured to the fascia.The locking anchor 172 uses a set screw 176 for locking purposes, and abi-directional disposable torque wrench for locking and unlocking.Tactile and audible feedback is provided for both locking and unlocking.

Both anchors 171, 172 can be molded from implant-grade silicone, but thelocking anchor 172 uses an internal titanium assembly for locking. The3-part mechanism is made of a housing 175, a locking set screw 176, anda blocking set screw 177 to prevent the locking set screw from back out.All three components can be titanium (6Al4V). The bi-directional torquewrench can have a plastic body and stainless steel hex shaft.

Wireless Dongle

The wireless dongle 200 is the hardware connection to asmartphone/mobile 202 or tablet 204 that allows communication betweenthe trial generator 107 or IPG 102 and the smartphone/mobile device 202or tablet 204, as illustrated in FIG. 14. During the trial or permanentimplant phases, the wireless dongle 200 is connected to the tablet 204through the tablet 204 specific connection pins and the clinicianprogrammer software on the tablet 204 is used to control the stimulationparameters. The commands from the clinician programmer software aretransferred to the wireless dongle 200 which is then transferred fromthe wireless dongle 200 using RF signals to the trial generator 107 orthe IPG 102. Once the parameters on the clinician programmers have beenset, the parameters are saved on the tablet 204 and can be transferredto the patient programmer software on the smartphone/mobile device 202.The wireless dongle 200 is composed of an antenna, a microcontroller(having the same specifications as the IPG 102 and trial generator 107),and a pin connector to connect with the smartphone/mobile device 202 andthe tablet 204.

Charger

The IPG 102 has a rechargeable lithium ion battery 108 to power itsactivities. An external induction type charger 210 (FIG. 1) wirelesslyrecharges the included battery 108 inside the IPG 102. The charger 210is packaged into a housing and consists of a rechargeable battery, aprimary coil of wire and a printed circuit board (PCB) containing theelectronics. In operation, charger 210 produces a magnetic field andinduces voltage into the secondary coil 109 in the IPG 102. The inducedvoltage is then rectified and used to charge the battery 108 inside theIPG 102. To maximize the coupling between the coils, both internal andexternal coils are combined with capacitors to make them resonate at aparticular common frequency. The coil acting as an inductor L forms anLC resonance tank. The charger uses a Class-E amplifier topology toproduce the alternating current in the primary coil around the resonantfrequency. The charger 210 features include, but are not limited to:

-   -   Charge IPG 102 wirelessly    -   Charge up to a maximum depth of 30 mm    -   Integrated alignment sensor indicates alignment between the        charger and IPG 102 resulting in higher power transfer        efficiency    -   Alignment sensor provides audible and visual feedback to the        user    -   Compact and Portable

A protected type of cylindrical Li ion battery is used as the charger210 battery. A Class-E power amplifier topology is a much used type ofamplifier for induction chargers, especially for implantable electronicmedical devices. Due to the Class-E power amplifier's relatively hightheoretical efficiency it is often used for devices where highefficiency power transfer is necessary. A 0.1 ohm high wattage resistoris used in series to sense the current through this circuit.

The primary coil L1 is made by 60 turns of Litz wire type 100/44-100strands of 44 AWG each. The Litz wire solves the problem of skin effectand keeps its impedance low at high frequencies. Inductance of this coilwas initially set at 181 uH, but backing it with a Ferrite plateincreases the inductance to 229.7 uH. The attached ferrite plate focusesthe produced magnetic field towards the direction of the implant. Such asetup helps the secondary coil receive more magnetic fields and aids itto induce higher power.

When the switch is ON, the resonance is at frequency

$f = \frac{1}{2\pi\sqrt{L\; 1C\; 2}}$When the switch is OFF, it shifts to

$f = \frac{1}{2\pi\sqrt{L\; 1\frac{C\; 1C\; 2}{{C\; 1} + {C\; 2}}}}$In a continuous operation the resonance frequency will be in the range

$\frac{1}{2\pi\sqrt{L\; 1C\; 2}} < f < \frac{1}{2\pi\sqrt{L\; 1\frac{C\; 1C\; 2}{{C\; 1} + {C\; 2}}}}$

To make the ON and OFF resonance frequencies closer, a relatively largervalue of C1 can be chosen by a simple criteria as follows

C1=nC2; a value of n=4 was used in the example above; in most cases3<n<10.

The voltages in these Class-E amplifiers typically go up to the order of300 VAC. Capacitors selected must be able to withstand these highvoltages, sustain high currents and still maintain low Effective SeriesResistance (ESR). Higher ESRs result in unnecessary power losses in theform of heat. The circuit is connected to the battery through aninductor which acts as a choke. The choke helps to smoothen the supplyto the circuit. The N Channel MOSFET acts as a switch in this Class-Epower amplifier. A FET with low ON resistance and with high draincurrent Id is desirable.

In summary, the circuit is able to recharge the IPG 102 battery 108 from0 to 100% in approximately two hours forty-five minutes with distancebetween the coils being 29 mm. The primary coil and the Class-Eamplifier draws DC current of 0.866 A to achieve this task. To improvethe efficiency of the circuit, a feedback closed loop control isimplemented to reduce the losses. The losses are minimum when the MOSFETis switched ON and when the voltage on its drain side is close to zero.

The controller takes the outputs from operational amplifiers, checks ifthe outputs meet the criteria, then triggers the driver to switch ON theMOSFET for the next cycle. The controller can use a delay timer, an ORgate and a 555 timer in monostable configuration to condition the signalfor driver. When the device is switched ON, the circuit will notfunction right away as there is no active feedback loop. The feedbackbecomes active when the circuit starts to function. To provide an activefeedback loop, an initial external trigger is applied to jump start thesystem.

Alignment Sensor

In some embodiments, the efficiency of the power transfer between theexternal charger 210 and the internal IPG 102 will be maximum only whenthe charger 210 and IPG 102 are properly aligned. An alignment sensorcan be used to provide proper alignment as part of the external circuitdesign and is based on the principle of reflected impedance. When theexternal coil is brought closer to the internal coil, the impedance ofboth circuits change. The sensing is based on measuring the reflectedimpedance and testing whether it crosses the threshold. A beeperprovides an audible feedback to the patient and a LED provides visualfeedback.

When the impedance of the circuit changes, the current passing throughit also changes. A high power 0.1 ohm resistor can be used in the seriesof the circuit to monitor the change in current. The voltage drop acrossthe resistor is amplified 40 times and then compared to a fixedthreshold value using an operational amplifier voltage comparator. Theoutput is fed to a timer chip which in turn activates the beeper and LEDto provide feedback to the user.

In some embodiments, the circuit can sense the alignment up to adistance of approximately 30 mm. In other embodiments, the circuit cansense the alignment up to a distance of approximately 20 mm, 40 mm or 50mm. The current fluctuation in the circuit depends on more factors thanreflected impedance alone and the circuit is sensitive to otherparameters of the circuit as well. To reduce the sensitivity related toother parameters, one option is to eliminate interference of all theother factors and improve the functionality of the reflected impedancesensor-which is very challenging to implement within the limited spaceavailable for circuitry. Another option is to use a dedicated sensorchip to measure the reflected impedance.

A second design uses sensors designed for proximity detector or metaldetectors for alignment sensing. Chips designed to detect metal bodiesby the effect of Eddy currents on the HF losses of a coil can be usedfor this application. The TDE0160 is an example of such a chip.

In some embodiments, the external charger is designed to work at 75 to80 kHz, whereas the proximity sensor was designed for 1 MHz. In otherembodiments, the external charger is designed to work at 60 to 90 kHz,whereas the proximity sensor is designed for 0.5-2 Mhz. The sensorcircuit is designed to be compatible with the rest of the external andis fine tuned to detect the internal IPG 102 from a distance ofapproximately 30 mm according to some embodiments.

Programmer

The Clinician Programmer is an application that is installed on a tablet204. It is used by the clinician to set the stimulation parameters onthe trial generator 107 or IPG 102 during trial and permanentimplantation in the operating room. The clinician programmer is capableof saving multiple settings for multiple patients and can be used toadjust the stimulation parameters outside of the operations room. It iscapable of changing the stimulation parameters though the RF wirelessdongle 200 when the trial generator 107 or IPG 102 which has beenimplanted in the patient is within the RF range. In addition, it is alsocapable of setting or changing the stimulation parameters on the trialgenerator 107 and/or the IPG 102 through the internet when both thetablet 204 and the Patient Programmers on a smartphone/mobile device 202both have access to the internet.

The Patient Programmer is an application that is installed on asmartphone/mobile device 202. It is used by the patient to set thestimulation parameters on the trial generator 107 or IPG 102 after trialand permanent implantation outside the operating room. The clinicianprogrammer is capable of saving multiple settings for multiple patientsand can be transferred to the Patient Programmer wirelessly when theClinician Programmer tablet 204 and the Patient Programmersmartphone/mobile device 202 are within wireless range such as Bluetoothfrom each other. In the scenario where the Clinician Programmer tablet204 and the Patient Programmer smartphone/mobile device 202 are out ofwireless range from each other, the data can be transferred through theinternet where both devices 202, 204 have wireless access such as Wi-Fi.The Patient Programmer is capable of changing the stimulation parameterson the trial generator 107 or IPG 102 though the RF wireless dongle 200when the trial generator 107 or IPG implanted in the patient is withinthe RF range.

Tuohy Needle

The Tuohy needle 240, as depicted in FIG. 15, is used in conjunctionwith a saline-loaded syringe for loss-of-resistance needle placement,and percutaneous stimulation leads 140, for lead 140 placement into thespinal canal. The Tuohy epidural needle 240 is inserted slowly into thespinal canal using a loss-of-resistance technique to gauge needle 240depth. Once inserted to the appropriate depth, the percutaneousstimulation lead 140 is passed through the needle 240 and into thespinal canal.

The epidural needle 240 is a non-coring 14G stainless steel spinalneedle 240 and will be available in lengths of 5″ (127 mm) and 6″(152.4). The distal tip 242 of the needle 240 has a slight curve todirect the stimulation lead 140 into the spinal canal. The proximal end246 is a standard Leur-Lock connection 248.

Stylet

The stylet 250, as depicted in FIG. 16, is used to drive the tip of apercutaneous stimulation lead 140 to the desired stimulation zone byadding rigidity and steerability. The stylet 250 wire 252 passes throughthe center lumen 142 of the percutaneous lead 140 and stops at theblocking plug at the distal tip of the lead 140. The tip of the stylet250 comes with both straight and curved tips. A small handle 254 is usedat the proximal end of the stylet 250 to rotate the stylet 250 withinthe center lumen 142 to assist with driving. This handle 254 can beremoved and reattached allowing anchors 170 to pass over the lead 140while the stylet 250 is still in place. The stylet 250 wire 252 is aPTFE coated stainless steel wire and the handle 254 is plastic.

Passing Elevator

The passing elevator 260, as depicted in FIG. 17, is used prior topaddle lead 141 placement to clear out tissue in the spinal canal andhelp the surgeon size the lead to the anatomy. The passing elevator 260provides a flexible paddle-shaped tip 262 to clear the spinal canal ofobstructions. The flexible tip is attached to a surgical handle 264.

The passing elevator 260 is a one-piece disposable plastic instrumentmade of a flexible high strength material with high lubricity. Theflexibility advantageously allows the instrument to easily conform tothe angle of the spinal canal and the lubricity allows the instrument toeasily pass through tissue.

Tunneling Tool

The tunneling tool 270, as depicted in FIG. 18, is used to provide asubcutaneous canal to pass stimulation leads 140 from the entrance pointinto the spinal canal to the IPG implantation site. The tunneling tool270 is a long skewer-shaped tool with a ringlet handle 272 at theproximal end 274. The tool 270 is covered by a plastic sheath 276 with atapered tip 278 which allows the tool 270 to easily pass through tissue.Once the IPG 102 implantation zone is bridge to the lead 140 entrancepoint into the spinal canal, the inner core 275 is removed, leaving thesheath 276 behind. The leads 140 can then be passed through the sheath276 to the IPG 102 implantation site. The tunneling tool 270 is oftenbent to assist in steering through the tissue.

The tunneling tool 270 is made of a 304 stainless steel core with afluorinated ethylene propylene (FEP) sheath 276. The 304 stainless steelis used for its strength and ductility during bending, and the sheath276 is used for its strength and lubricity.

Torque Wrench

The torque wrench 280, as depicted in FIG. 19, is used in conjunctionwith the IPG 102, lead extension 150 and lead splitter 160 to tightenthe internal set screw 119, which provides a radial force against thefixation contact of the stimulation leads 140, 141, preventing the leads140, 141 from detaching. The torque wrench 280 is also used to lock andunlock the anchor 170. The torque wrench 280 is a small, disposable,medical instrument that is used in every SCS 100 case. The torque wrench280 advantageously provides audible and tactile feedback to the surgeonthat the lead 140, 141 is secured to the IPG 102, extension 150, orsplitter 160, or that the anchor 170 is in the locked or unlockedposition.

The torque wrench 280 is a 0.9 mm stainless steel hex shaft 282assembled with a plastic body 284. The wrench's 280 torque rating isbi-directional, primarily to provide feedback that the anchor 170 iseither locked or unlocked. The torque rating allows firm fixation of theset screws 119, 152 against the stimulation leads 140, 141 withoutover-tightening.

Trial Patch

The trial patch is used in conjunction with the trialing pulse generator107 to provide a clean, ergonomic protective cover of the stimulationlead 140, 141 entrance point in the spinal canal. The patch is alsointended to cover and contain the trial generator 107. The patch is alarge, adhesive bandage that is applied to the patient post-operativelyduring the trialing stage. The patch completely covers the leads 140,141 and generator 107, and fixates to the patient with anti-microbialadhesive.

In some embodiments, the patch is a watertight, 150 mm×250 mmanti-microbial adhesive patch. The watertight patch allows patients toshower during the trialing period, and the anti-microbial adhesivedecreases the risk of infection. The patch will be made of polyethylene,silicone, urethane, acrylate, and rayon.

Magnetic Switch

The magnetic switch is a magnet the size of a coin that, when placednear the IPG 102, can switch it on or off. The direction the magnet isfacing the IPG 102 determines if the magnetic switch is switching theIPG 102 on or off.

As shown in FIG. 20, the implantable pulse generator (IPG) 102 includesan RF transceiver module 103, a processor such as the microcontroller104 and a programmable signal generator such as the ASIC 106. Thetransceiver module 103 manages wireless communication between themicrocontroller 104 and external remote (e.g., dongle 200 connected toeither the smartphone/mobile 202 or tablet 204).

In the embodiment shown in FIG. 20, a treatment control module 14 storedin a flash memory 13 of the microcontroller 104 is executed by themicrocontroller to centrally control operation of every component andcircuits of the IPG 102 with the exception of an independently operatedcharger (not shown) that charges the battery 108. Specifically, thetreatment control module 14 handles programming of the RF transceivermodule 103 and signal generator 106 among other functions.Communications among the microcontroller 104, RF transceiver module 103and signal generator 106 are performed over a bus 16 such as the SerialPeripheral Interface (SPI) bus.

One exemplary microcontroller 104 may be MSP430F5328 from TexasInstruments of Dallas, Tex. as it has very low power usage, large amountof memory, integrated peripherals and small physical size.

As shown in more detail in FIG. 21, the signal generator 106 includesmemory (control registers 18), timing generator 20, arbitrator circuit22, high frequency generator 24, electrode driver 26 which are allcoupled to each other. All components in FIG. 21 have access to and aresupplied with signal parameters stored in the control registers 18through a register bus 28.

One of the many novel features of the IPG 102 is that the controlregisters 18 in the signal generator 106 have sufficient memory to storeall of the signal parameters necessary to drive the electrodes E1-E32independently of the microcontroller 104. As a result, themicrocontroller 104 can be placed in a standby mode once it programs allof the pulse parameters in the control registers 18 and the treatmentcontrol module 14 instructs the signal generator to generate thestimulation signals by setting the stimulation enable pin STIM-EN tologic high. In the embodiment shown in which the microcontroller isMSP430F5328, the treatment control module 14 places the microcontrollerin LPM3 Standby Mode. In an LPM3 mode, the master clock (main clock)that drives the instruction execution unit of the microcontroller 104 isturned off, essentially turning the microcontroller off so as toconserve battery power.

The microcontroller can be waken up from the standby mode by aninterrupt signal IRQ which can be transmitted by the transceiver module103 when it receives an appropriate instruction from a remote controldevice.

As discussed earlier, in the IPG 102, the electrodes E1-E32 may begrouped into stimulation sets (stimulation channels). Each stimulationchannel represents one particular stimulation signal/pattern which isapplied to the associated electrodes. In the embodiment shown, the IPG102 can accommodate up to 16 channels (ch1 through ch16). In someembodiments, each electrode can belong to one or more channels up to themaximum number of channels and each channel can be associated with atleast 2 electrodes to a maximum of 32 electrodes. Accordingly, oneelectrode can belong to all 16 channels. In some embodiments, eachchannel can be associated with more than 32 electrodes, such as 40electrodes.

For example, as shown in FIG. 25, channel 1 includes electrodes E1 andE2, channel 2 includes electrodes E2 and E3 while channel 3 includeselectrodes E2-E4 and E6-E8. Thus, electrode E2 belongs to channels 1, 2and 3, electrode E3 belongs to channels 2 and 3, while electrodes E4 andE5-E8 belong to only channel 3 and electrode E1 belong to only channel1. In FIG. 25, electrode E5 is unused and therefore does not belong toany channel. Data that associates electrodes to particular stimulationchannels are stored in the control registers 18.

As will be discussed more fully later herein, for each channel, the IPG102 is capable of programming the amplitude, frequency and duration ofboth the first phase pulse (pulse1) and the second phase pulse (pulse2)of a biphasic pulse (see FIG. 26, for example) for all of thestimulation channels. As seen in FIG. 26, waveforms of a biphasic pulsecomprising pulse1 and pulse2 of an active channel (ch_act) are combinedinto a single pulse pattern which are applied to the electrodesbelonging to the active channel such as E1 and E2. FIG. 26 illustratesthe pulse width and amplitude variations. As can be seen, the width ofpulse1 is wider than that of pulse2 while the amplitude for pulse1 asscaled by the pulse scaler 38 (see positive pulse of E1, for example) islower than that of pulse2 (see negative pulse of E1, for example).

Advantageously, the pulse parameters for pulse1 are independent of thosefor pulse2 for maximum flexibility in managing pain. As an example andreferring to FIG. 25, assume that the current path between electrodes E1and E2 (channel 1) affects a nerve path to a left leg while the currentpath between electrodes E2 and E3 (channel 2) affects a nerve path to aright leg. When the patient complains of more pain on the right leg thanthe left leg, a physician programming the IPG 102 may associateelectrodes E1 and E2 to channel 1 with a stimulation pulse patternhaving 100 Hz in frequency and associate electrodes E2 and E3 to channel2 with a pulse pattern having 1000 Hz in frequency and a higher currentamplitude than channel 1. In this way, smaller current is applied tochannel 1 for the left leg and higher current is applied to channel 2for the right leg. Advantageously, the flexibility of the IPG 102 allowsjust the right amount of current to each affected area of the patient.

The control registers 18 include standard read/write registers 18 thatcan be accessed by the SPI bus 16 and register bus 28. The controlregisters 18 are configured as an array of 8-bit registers, each with aunique address. The control registers 18 are programmed by themicrocontroller 104 to store all pulse parameters that are necessary forthe signal generator 106 to generate all of the stimulation channelpatterns without any intervention from the microcontroller. The pulseparameters include stimulation channel timing settings, current (pulseamplitude) scaler settings, calibration data and electrode groupparameters.

For each channel, the control registers 18 store the rising and fallingedges of the channel itself, rising and falling edges of each of the twopulses (pulse1 and pulse2), period of the biphasic pulse, active channelperiod (channel envelope), and current scaling (pulse amplitude) valuesfor both pulses (pulse1 and pulse2). For each channel, the controlregisters 18 also store burst frequency data (as will be explained laterherein) such as burst period for both pulses (pulse1 and pulse2). Foreach channel, the control registers 18 also store data regarding whichelectrodes E1-E32 belong to that channel. For each channel and for eachelectrode within that channel, the control registers 18 storesource/sink data for both pulses of the biphasic pulse (pulse1 andpulse2), i.e., whether each electrode will be sourcing current orsinking current during pulse1 and pulse2. All parameters are specifiedwith reference to the origin and is in units of microseconds.

The timing generator 20 generates stimulation timing signals whichcomprise pulse1, pulse2, and a channel pulse “ch” (channel envelopewaveform) for all 16 channels based on the pulse timing parametersstored in the control registers 18. If all 16 channels are programmed bythe clinician, then the timing generator 20 generates the pulse1,pulse2, and channel pulse data ch for all 16 channels simultaneously.

As an example, pulse1 and pulse2 waveforms of FIG. 26 illustrate theoutput waveforms from the timing generator 20 for an exemplary biphasicpulse of a particular channel. The channel envelope data ch_act definesthe start and end of an active portion of each channel as well as thechannel period, which can be defined as the time between two adjacentrising edges of the channel. Pulse1 and Pulse2 define the start and endof each of the two phases of the biphasic pulse.

Because of the flexibility of the IPG 102, more than one channel couldbecome active at any time when stimulation patterns of multiple channelsare programmed. The arbitrator 22 is designed to resolve the overlappingchannel (channel contention) problem by ensuring that only one channelis active at any one time. Among others, the circuits in the arbitrator22 are designed with two rules. The first rule is that when an activechannel is being selected (i.e., a channel currently in progress), allother channels attempting to go active are suppressed and discarded. Thesecond rule is that when two or more channels are about to become activewith simultaneous rising edges in ch, an active channel will bedetermined based on a predetermined channel priority.

In the embodiment shown, the arbitrator 22 has been programmed such thatthe lowest numbered channel will be given priority and the remainingsimultaneous channels will be discarded. Since there are 16 channels(ch1 through ch16) in the IPG 102, channel one has the highest prioritywhile channel 16 has the lowest priority.

The output of the arbitrator 22 includes pulse timing signal p1_act andp2_act which are the same waveforms as pulse1 and pulse2 of an activechannel. The arbitrator 22 also outputs the channel envelope of anactive channel (ch_act as shown in FIG. 26, for example) for use by thehigh frequency generator 24 as well as the channel number of the activechannel (ch_code), which will be used by the electrode driver 26, aswill be explained later herein. In the embodiment shown, the channelnumber is a 4-bit code that identifies the number of the active channel.For example, ‘0001’ represents channel 2 while ‘1111’ represents channel16.

FIG. 27 provides an example of the channel arbitration by the arbitrator22. In FIG. 27, channel number one has 3 electrodes E1-E3 and channelnumber two has 2 electrodes E1-E2. As the channel period for the twochannels is different, they will overlap from time to time. In theillustration, channel one is active (ch1_active) when channel twoattempts to become active. At that time, the arbitrator executes thefirst rule, and will suppress channel two and prevent it from becomingactive. Thus, only the pulse1/pulse2 signals that drive the electrodesfor channel one (active channel) will be output by the arbitrator 22.The pulse1/pulse2 signals and channel envelope signal for channel two(ch2_active) are shown in dotted lines to show that they have beensuppressed by the arbitrator 22.

The high frequency generator 24 receives the p1_act and p2_act waveformsfrom the arbitrator 22, decides whether to modulate the received signalsbased on the stored parameters in the control registers 18. If thedecision is no, then the high frequency generator 24 passes the receivedpulse signals unaltered to the electrode driver 26.

If the decision is a yes, however, the high frequency generator 24modulates the received signals at a burst frequency that has beenprogrammed into the control registers 18. The burst frequency is higherthan the frequency of the received signals p1_act and p2_act.

The electrode driver 26 receives the output (p1, p2 and ch_code) of thehigh frequency generator 24, amplifies the received signal according tothe pulse amplitude parameters stored in the control registers 18, andoutputs the final stimulation pattern for each channel to be appliedthrough the electrodes E1-E32. As discussed above, the burst pulseparameters stored in the control registers 18 have separate frequencyvalues for pulse1 and pulse2 such that an asymmetric pulse shape withpositive and negative pulses having different frequency values can begenerated by the electrode driver 26.

FIG. 22 is a more detailed functional block diagram of the highfrequency generator of FIG. 21. The high frequency generator 24 includesa burst generator 30 and burst multiplexer 32.

The burst multiplexer 32 receives burst parameters stored in the controlregisters 18 for all the channels, selects the burst parametersassociated with an active channel, and outputs the selected burstparameters to the burst generator 30. Specifically, there is apulse1/pulse2 burst register pair for each channel for the burst option,totaling 32 registers for the 16 possible channels in the embodimentshown. The burst multiplexer 32 is a vector MUX that selects thepulse1/pulse2 register pair corresponding to the active channel number.The select lines 34 to the burst multiplexer 32 is the active channelnumber (ch_code) from the arbitrator 22, which identifies the activechannel at any given time. The selected burst parameters forpulse1/pulse2 are sent to the burst generator 30.

Within the burst generator 30, there are 2 independent burst generatorcircuits, one for pulse1 and one for pulse2. The pulse1 and pulse2signals belonging to the active channel are passed through to the burstgenerator circuits from the arbitrator 22. If the pulse1/pulse2 burstparameter data stored in the control registers 18 is zero (therefore theselected burst parameters to the burst generator 30 are also zero), thenno burst is generated, in which case the pulse1/pulse2 signals from thearbitrator 22 are sent unaltered to the electrode driver 26. If thepulse1/pulse2 burst register in the control registers 18 is programmed,then the pulse1/pulse2 duration will be replaced by (modulated to) thecorresponding programmed burst signal based on the selected burstparameters from the burst multiplexer 32.

As an example, FIG. 27 illustrates that pulse1 for channel one has beenprogrammed for high frequency modulation while pulse2 for the samechannel has not been programmed. Specifically, a single pulse1 pulse hasbeen modulated to (replaced with) five higher frequency burst pulses.Thus, the frequency of the newly modulated pulse1 is five times thefrequency of the original pulse1 signal.

FIG. 23 is a more detailed functional block diagram of the electrodedriver 26 of FIG. 21. The electrode driver 26 includes a pulse amplitudemultiplexer 36, pulse scaler 38 and current drivers 40.

The pulse amplitude multiplexer 36 receives amplitude parameters storedin the control registers 18 for all the channels, selects the amplitudeparameters associated with an active channel, and outputs the selectedamplitude parameters to the pulse scaler 38.

Specifically, 512 bytes (16 by 32 bytes-32 bytes for each channel) inthe control registers 18 are reserved for storing pulse amplitude data.Each of the 16 channels is associated with 32 bytes with each byterepresenting pulse amplitude information for pulse1 and pulse2 of eachof the 32 electrodes E1-E32. From one byte, 7 bits are used to store theamplitude information for pulse1 and pulse2 and the remaining bit (MSB)defines the polarity of the pulse at the associated electrode as will bediscussed later herein.

Similar to the burst multiplexer 32, the pulse amplitude multiplexer 36is a vector MUX that selects the 32 bytes of amplitude parameters forpulse1/pulse2 corresponding to the active channel number. The selectlines 34 to the amplitude multiplexer 36 is the active channel number(ch_code) from the arbitrator 22, which identifies the active channel atany given time. The selected amplitude parameters of pulse1/pulse2 forall 32 electrodes E1-E32 are sent to the pulse scaler 38. The pulsescaler 38 outputs amplitude scaling factors for all electrodes of theactive channel. Thus, the pulse scaler 38 includes 32 identical scalerscorresponding to the 32 electrodes E1-E32. In the embodiment shown, eachscaler includes a D/A converter that converts the digital amplitudevalue into a corresponding analog value 11-132.

As discussed above, the signal generator 106 supports asymmetrical pulseamplitude feature, which means the amplitude for pulse1 and pulse2 canbe different. The amplitude scaling data are stored in the controlregisters 18. In the embodiment shown, 4 bits are used to specify thescaling factor for pulse1 and pulse2 for each electrode-2 bits forpulse1 and 2 bits for pulse2. Moreover, the signal generator 106 cansupport asymmetrical pulse width variation between pulse1 and pulse2.

The pulse scaler 38 adjusts the amplitude parameter by the associatedscaling factor stored in the associated 2 bits for pulse1 and pulse2. Inthe embodiment shown, the pulse scaler 38 performs the scaling functionby shifting to the right the content of the selected amplitude parameter(7 bits of data from the amplitude multiplexer 36) by the number storedin the corresponding 2 bit scaling factor. Since there are fourpossibilities in a 2 bit number (0, 1, 2, or 3), the amplitude can bereduced by ½, ¼ or ⅛.

For example, assume that the active channel is channel one, theamplitude parameter for electrode E1 for channel one in the controlregisters 18 is binary “1111111” (decimal 127) while the associated 2bit scaling factor is binary “11” for pulse1 and “01” for pulse2. Thenumber 127 represents 12.7 mA of current. For pulse1, the 7 bitamplitude content will be shifted to the right by 3 bits (binary “11”)for pulse1 and by 1 bit (binary “01”) for pulse2. Thus, the pulse scaler38 will scale the amplitude of 127 down to 15 (binary “1111”) for pulse1and to 63 (binary “111111”) for pulse2, which corresponds to a currentof 1.5 mA for pulse1 and 6.3 mA for pulse2.

In the embodiment shown, the scaling factor parameters are stored inunused bits of burst parameters and are passed to the pulse scaler 38from the burst multiplexer 32. However, dedicated memory can beallocated in the control registers 18. It is to be noted that thestimulation programming software should ensure that at any given instantof time, the algebraic sum of all electrode currents is zero, and the dcaverage of the current per cycle at each electrode is also zero.

Another function performed by the pulse scaler 38 is that it convertsthe scaled amplitude digital value into an analog output current that islinearly proportional to the value of the digital value. The analogoutput current for each electrode is supplied to the current drivers 40.In the embodiment shown, the analog output current represents 1/20 ofthe actual current to be supplied to the associated electrode.

The current drivers 40 amplify the analog output currents from the pulsescaler 38 and switches the amplified current to the appropriateelectrodes E1-E32 based on the pulse parameters stored in the controlregisters 18. In some embodiments, when there are 32 electrodes, thereare 32 electrode drivers 40 in the embodiment shown. In the embodimentshown, each current driver 40 is a current driver that amplifies theinput signal from the pulse scaler 38 by 20 times.

FIG. 24 is a functional illustration of two of the current drivers ofFIG. 23. Each electrode driver 40 can be a current source that can sinkor source current whose amplitude is based on a control current signalcoming from the pulse scaler 38. In the embodiment shown, the currentsource includes a pair of NMOS current source 42 and PMOS current source44 that are coupled in series between the voltage supply and ground.Each of the PMOS and NMOS current sources can be implemented as acurrent mirror in a well-known manner. The NMOS current source 44 sinkscurrent from the associated electrode to ground while the PMOS currentsource 44 sources current from the positive voltage supply to theassociated electrode. Each electrode driver 40 includes switches 46 and48 connected in series between the voltage supply and ground to eithersource or sink the current.

The control registers 18 store the pulse parameter that relate towhether a particular electrode will be sourcing current or sinkingcurrent. In the embodiment shown, bit 7 (MSB) of each byte of the pulseamplitude parameters that are supplied to the burst multiplexer 32 isused to specify whether a particular current source 40 will be sourcingcurrent or sinking current. If the bit is zero, during pulse1, switch 46will be turned on while switch 48 will be turned off, and during pulse2,switch 46 will be turned off while switch 48 will be turned on. If thebit is set (i.e., it is a “1”), during pulse1, switch 46 will be turnedoff while switch 48 will be turned on, and during pulse2, switch 46 willbe turned on while switch 48 will be turned off.

FIG. 24 illustrates the current path for two electrodes Ej and Ek whenbit 7 for Ej=0 and bit 7 for Ek=1. During pulse1, switch 46 for Ej andswitch 48 for Ek turn on to create a current path 50. In that instance,the PMOS current source 42 for Ej will be sourcing current while theNMOS current source 44 for Ek will be sinking current to ground.Conversely, during pulse2, switch 46 for Ek and switch 48 for Ej turn onto create a current path 52. In that instance, the PMOS current source42 for Ek will be sourcing current while the NMOS current source 44 forEj will be sinking current to ground.

The pulse shape at the electrode Ej will look similar to the E1 waveformas shown in FIG. 26 while the pulse shape at the electrode Ek will looksimilar to the E2 waveform.

The foregoing specific embodiments represent just some of the ways ofpracticing the present invention. Many other embodiments are possiblewithin the spirit of the invention. Accordingly, the scope of theinvention is not limited to the foregoing specification, but instead isgiven by the appended claims along with their full range of equivalents.

What is claimed is:
 1. A system for spinal cord stimulation comprising:a plurality of stimulation electrodes, wherein each of the stimulationelectrodes are used to provide electrical pulse therapy to a spinal cordof a patient; a plurality of leads, wherein each of the plurality ofleads extends to one of the stimulation electrodes; a lead anchorimplantable into the patient, the lead anchor configured to lock one ofthe plurality of leads to tissue of the patient to prevent migration ofa distal tip of the one of the plurality of leads, wherein the leadanchor includes a locking set screw; an implantable pulse generatorconfigured to be implantable into the patient and connectable to theplurality of leads, wherein the implantable pulse generator isconfigured to generate signals to transmit to the plurality ofstimulation electrodes; and a trial generator configured to betemporarily connectable to the plurality of leads to determine thesignals generated by the implantable pulse generator wherein the leadanchor includes a blocking set screw, positioned in the lead anchor, toprevent the locking set screw from backing out.
 2. The system of claim1, wherein the implantable pulse generator comprises a transceiver andan ASIC.
 3. The system of claim 2, wherein the ASIC comprises a digitalsection and an analog section.
 4. The system of claim 3, wherein thedigital section comprises registers configured to store informationrelating to the electrical pulse therapy.
 5. The system of claim 1,wherein the implantable pulse generator is recharged wirelessly via aninduction coil.
 6. The system of claim 1, wherein the plurality of leadscomprise percutaneous stimulation leads.
 7. The system of claim 6,wherein the plurality of leads are accompanied by a suture configured tosecure the one or more leads around tissue.
 8. The system of claim 6,wherein the percutaneous stimulation leads are flexible.
 9. The systemof claim 8, wherein the plurality of leads are paddle stimulation leads.10. The system of claim 1, wherein the plurality of leads are configuredto drive into a spinal canal of the patient by using a steering stylet.11. A system for spinal cord stimulation comprising: one or morestimulation electrodes, wherein the one or more stimulation electrodesare used to provide electrical pulse therapy to a spinal cord of apatient; one or more leads extending to the one or more stimulationelectrodes; at least one lead anchor implantable into the patient andconfigured to lock the one or more leads to tissue of the patient toprevent migration of a distal tip of each of the one more leads, whereinthe at least one lead anchor includes a locking set screw; animplantable pulse generator configured to be implantable into thepatient and connectable to the one or more leads, wherein theimplantable pulse generator is configured to generate signals totransmit to the one or more stimulation electrodes; a charger forcharging the implantable pulse generator wirelessly; and a trialgenerator configured to be temporarily connectable to the one or moreleads to determine the signals generated by the implantable pulsegenerator, wherein the at least one lead anchor includes a blocking setscrew, positioned in the at least one lead anchor, to prevent thelocking set screw from backing out.
 12. The system of claim 11, whereinthe implantable pulse generator comprises a transceiver and an ASIC. 13.The system of claim 12, wherein the ASIC comprises a digital section andan analog section.
 14. The system of claim 13, wherein the digitalsection comprises registers configured to store information relating tothe electrical pulse therapy.
 15. The system of claim 11, wherein theimplantable pulse generator is recharged wirelessly via an inductioncoil.
 16. The system of claim 11, wherein the one or more leads comprisepercutaneous stimulation leads.
 17. The system of claim 16, wherein theone or more leads are accompanied by a suture configured to secure theone or more leads around tissue.
 18. The system of claim 16, wherein thepercutaneous stimulation leads are flexible.
 19. The system of claim 18,wherein the one or more leads are paddle stimulation leads.
 20. Thesystem of claim 11, wherein the one or more leads are configured todrive into a spinal canal of the patient by using a steering stylet.